System and method for delta relaxation enhanced magnetic resonance imaging

ABSTRACT

A delta-relaxation magnetic resonance imaging (DREMR) system is provided. The system includes a main field magnet and field shifting coils. A main magnetic field with a strength B0 can be generated using the main filed magnet and the strength B0 of the main magnetic field can be varied through the use of the field-shifting coils. The DREMR system can be used to perform signal acquisition based on a pulse sequence for acquiring at least one of T2*-weighted signals imaging; MR spectroscopy signals; saturation imaging signals and MR signals for fingerprinting. The MR signal acquisition can be augmented by varying the strength B0 of the main magnetic field for at least a portion of the pulse sequence used to acquire the MR signal.

CROSS REFERENCE TO RELATED APPLICATIONS

This is a continuation of U.S. patent application Ser. No. 14/902,221,filed Dec. 30, 2015, which is a 37 CFR 371 (c) national stage ofInternational Patent Application No. PCT/CA2015/000106, filed Feb. 23,2015, the contents of which are incorporated herein by reference.

FIELD OF THE INVENTION

The present invention relates generally to magnetic resonance imaging.More specifically, the present invention relates to delta relaxationenhanced magnetic resonance imaging.

BACKGROUND OF THE INVENTION

Magnetic resonance imaging (MRI) is a major imaging technique used inmedicine. MRI is capable of generating detailed images of soft tissuessuch as the brain, muscles and kidneys. Specific properties of thevarious compounds found inside tissues, such as water and/or fat, areused to generate images. When subjected to a strong magnetic field, thevector sum of the nuclear magnetic moments of a large number of atomspossessing a nuclear spin angular momentum, such as hydrogen, which isabundant in water and fat, will produce a net magnetic moment inalignment with the externally applied field. The resultant net magneticmoment can furthermore precess with a well-defined frequency that isproportional to the applied magnetic field. After excitation by radiofrequency pulses, the net magnetization will generate a signal that canbe detected.

Delta relaxation enhanced magnetic resonance imaging (DREMR) generallyreferred to as field-cycled relaxometry or field-cycled imaging is anMRI technique that offers the possibility of using underlying tissuecontrast mechanism which vary with the strength of the applied magneticfield to generate novel image contrasts. To achieve DREMR contrast, themain magnetic field is varied as a function of time during specificportions of an MR pulse sequence. A field-shifting electromagnet coil isused to perform the field variation. To date the DREMR imaging methodshave focused on the effect of main magnetic field variations on the T1relaxation characteristic of materials being imaged. This, however, is alimited use of a DREMR system.

SUMMARY OF THE INVENTION

It is an object to provide a novel system and method for an MRI scanningsystem and method that obviates and mitigates at least one of theabove-identified disadvantages of the prior art.

According to one aspect, a method of acquiring magnetic resonance (MR)signals at a delta-relaxation enhanced MR imaging (DREMR) system isprovided. According to the method, the DREMR system can generate a mainmagnetic field with a strength of B0 and an initial pulse sequence foracquiring at least one of: T2*-weighted MR imaging signals;susceptibility weighted imaging (SWI) signals; and saturation imagingsignals. The main magnetic field strength can be varied to a strength ofB1 during at least one portion of the initial pulse sequence and a firstimage can be acquired based on the initial pulse sequence.

According to another aspect, a method of acquiring MR signals at a DREMRsystem is provided. According to the method, the DREMR system cangenerate a main magnetic field with a strength of B0 and an initialpulse sequence for acquiring MR spectroscopy signals. A firstspectroscopy signal can be acquired based on the initial pulse sequence.A repeat pulse sequence for acquiring MR spectroscopy signals can alsobe generated, the repeat pulse sequence corresponding to the initialpulse sequence. The main magnetic field strength can be varied to astrength of B1 during at least one portion of the repeat pulse sequence.A second spectroscopy signal can be acquired based on the repeat pulsesequence and peaks from the first and the second spectroscopy signalscan be identified. The identified peaks can then be correlated.

According to yet another aspect, a method of acquiring MR signals at aDREMR system is provided. According to the method, the DREMR system cangenerate a main magnetic field with a strength of B0 and an initialpulse sequence for acquiring MR signals for fingerprinting. A firstimage can be acquired based on the initial pulse sequence. A repeatpulse sequence for acquiring MR fingerprinting signals can be generated,the repeat pulse sequence corresponding to the initial pulse sequence.The main magnetic field strength can be varied to a strength of B1during at least one portion of the repeat pulse sequence and a secondimage can be acquired based on the repeat pulse sequence. At least oneMR signal property can be measured based on the first and the secondimages. A tissue type can be identified based on the at least one MRsignal property.

According to a further aspect a DREMR system is provided. The system cancomprise a main magnet operating to generate a main magnetic field witha strength of B0. The system can further comprise radio frequency coilshaving a transmit aspect and gradient coils operating to generate aninitial pulse sequence for acquiring at least one of: T2*-weighted MRimaging signals; susceptibility weighted imaging (SWI) signals; andsaturation imaging signals. The system can also comprise field-shiftingmagnets operating to vary the main magnetic field strength to a strengthof B1 during at least one portion of the initial pulse sequence. Theradio frequency coils can have a receive aspect operating to acquire afirst image based on the initial pulse sequence.

These, together with other aspects and advantages which will besubsequently apparent, reside in the details of construction andoperation as more fully hereinafter described and claimed, referencebeing had to the accompanying drawings forming a part hereof, whereinlike numerals refer to like parts throughout.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 shows a block diagram of functional subsystems of a deltarelaxation magnetic resonance imaging (DREMR) system in accordance withan implementation;

FIG. 2 an imaging volume and corresponding slice to be scanned by thedelta relaxation magnetic resonance system of FIG. 1 in accordance withan implementation;

FIG. 3 shows illustrative examples of T1 and T2 relaxation diagrams;

FIG. 4 shows an example pulse sequence in accordance with animplementation;

FIG. 5 shows a schematic representation of a k-space containing onereceived line in accordance with an implementation;

FIG. 6 shows idealized frequency distribution of two materials atdifferent magnetic field strengths;

FIG. 7 shows an example pulse sequence for augmented MR signalacquisition using the example DREMR system of FIG. 1 based on spectralsuppression;

FIG. 8 shows an example pulse sequence for augmented MR signalacquisition using the example DREMR system of FIG. 1 based onsusceptibility weighted imaging;

FIG. 9 shows an example pulse sequence for augmented MR signalacquisition using the example DREMR system of FIG. 1 based onsusceptibility weighted imaging;

FIG. 10 shows an example pulse sequence for augmented MR signalacquisition using the example DREMR system of FIG. 1 based on T2* basedimaging;

FIG. 11 shows a conceptual illustration of T2* signal separation from 2materials; and

FIG. 12 shows an example pulse sequence for augmented MR signalacquisition using the example DREMR system of FIG. 1 based on T2* basedimaging;

FIG. 13 shows idealized results of performing augmented MR signalacquisition using the example DREMR system 100 of FIG. 1; and

FIG. 14 shows a simplified example of the effects of magnetic fieldstrength changes to MR fingerprinting results.

DETAILED DESCRIPTION

Referring to FIG. 1, a block diagram of a delta relaxation magneticresonance imaging (DREMR) system, in accordance with an exampleimplementation, is shown at 100. The example implementation of the DREMRsystem indicated at 100 is for illustrative purposes only, andvariations including additional, fewer and/or varied components arepossible. Traditional magnetic resonance imaging (MRI) systems representan imaging modality which is primarily used to construct pictures ofmagnetic resonance (MR) signals from protons such as hydrogen atoms inan object. In medical MRI, typical signals of interest are MR signalsfrom water and fat, the major hydrogen containing components of tissues.DREMR systems use field-shifting magnetic resonance methods inconjunction with traditional MRI techniques to obtain images withdifferent contrast than is possible with traditional MRI, includingmolecularly-specific contrast.

As shown in FIG. 1, the illustrative DREMR system 100 comprises a dataprocessing system 105. The data processing system 105 can generallyinclude one or more output devices such as a display, one or more inputdevices such as a keyboard and a mouse as well as one or more processorsconnected to a memory having volatile and persistent components. Thedata processing system 105 can further comprise one or more interfacesadapted for communication and data exchange with the hardware componentsof MRI system 100 used for performing a scan.

Continuing with FIG. 1, example the DREMR system 100 can also include amain field magnet 110. The main field magnet 110 can be implemented as apermanent, superconducting or a resistive magnet, for example. Othermagnet types, including hybrid magnets suitable for use in the DREMRsystem 100 will now occur to a person of skill and are contemplated. Themain field magnet 110 is operable to produce a substantially uniformmain magnetic field having a strength B0 and a direction along an axis.The main magnetic field is used to create an imaging volume within whichdesired atomic nuclei, such as the protons in Hydrogen within water andfat, of an object are magnetically aligned in preparation for a scan. Insome implementations, as in this example implementation, a main fieldcontrol unit 115 in communication with data processing system 105 can beused for controlling the operation of the main field magnet 110.

The DREMR system 100 can further include gradient coils 120 used forencoding spatial information in the main magnetic field along, forexample, three perpendicular gradient axis. The size and configurationof the gradient coils 120 can be such that they produce a controlled anduniform linear gradient. For example, three paired orthogonalcurrent-carrying primary coils located within the main field magnet 110can be designed to produce desired linear-gradient magnetic fields.

In some implementations, the gradient coils 120 may be shielded andinclude an outer layer of shield coils which can produce a countermagnetic field to counter the gradient magnetic field produced by theprimary gradient coils forming a primary-shield coils pair. In such acoil pair the “primary” coils can be responsible for creating thegradient field and the “shield” coils can be responsible for reducingthe stray field of the primary coil outside a certain volume such as animaging volume. The primary-shield coils pair of the gradient coils 120,the primary and shield coils, may be connected in series. It is alsopossible to have more than two layers of coils for any given gradientaxis that together form a shielded gradient coil. Shielded gradientcoils 120 may reduce eddy currents and other interference which cancause artefacts in the scanned images. Since eddy currents mainly flowin conducting components of the DREMR system 100 that are caused bymagnetic fields outside of the imaging volume (fringe fields), reducingthe fringe fields produced by the gradient coils 120 may reduceinterference. Accordingly, the shapes and sizes, conductor wire patternsand sizes, and current amplitudes and patterns of the primary-shieldcoils pair can be selected so that the net magnetic field outside thegradient coils 120 is as close to zero as possible. For cylindricalmagnets, for example, the two coils can be arranged in the form ofconcentric cylinders whereas for vertical field magnets, the two coilsmay be arranged in coaxial disks.

One side effect of shielding can be that the fields produced by theprimary-shield coils pair of the gradient coils 120 may partially canceleach other within the imaging volume. Accordingly, more current can berequired to produce a gradient field with a particular strength byshielded gradient coils 120 than by unshielded gradient coils 120. Thiseffect can be quantified as the gradient efficiency, which may bedefined as the achievable gradient strength for 1 Ampere of drivingcurrent. Another important parameter describing gradient coilperformance is called the gradient slew rate, which is the rate ofdriving a gradient coil from zero to its maximum amplitude. This term isinversely proportional to the inductance of the gradient coil.Typically, in order to increase the efficiency of a shielded gradientcoils 120 to be comparable to the efficiency of an unshielded gradientcoils 120 the inductance must increase. This increase in inductance willdecrease the maximum achievable slew rate. The loss in efficiency for ashielded configuration can depend on the distance and current densityratio between the primary and shield coils. Increasing the distancebetween the primary-shield coils pair may increase the efficiency.

The conductive components of the gradient coils 120, whether shielded orunshielded and including the primary and shield coils, may consist of anelectrical conductor (for example copper, aluminum, etc.). The internalelectrical connections can be such that when a voltage difference isapplied to the terminals of the gradient coils 120, electric current canflow in the desired path. The conductive components for the threegradient axes for both the primary gradient coils and the gradientshield coils can be insulated by physical separation and/or anon-conductive barrier. The primary gradient windings can be placed on anon-conductive substrate (for example, G10, FR4, epoxy or others).

In some variations, the gradient coils 120 may also be provided withthermal control or heat extraction mechanisms. For example, some of thewindings can be hollow and coolant can be passed through these hollowconductors to extract heat from the gradient coils 120, produced, forinstance, by resistive heating of the windings when electricity isapplied. Alternatively, other methods of extracting heat can be used,such as inserting coolant channels within the gradient coils 120. Thecoolant channels can be in thermal contact with the gradient coilwindings. The gradient coils 120 can also be mounted in athermally-conductive but electrically-non-conductive epoxy to ensurethat the mechanical assembly is rigid and to limit the possibility ofelectrical breakdown.

The magnetic fields produced by the gradient coils 120, in combinationand/or sequentially, can be superimposed on the main magnetic field suchthat selective spatial excitation of objects within the imaging volumecan occur. In addition to allowing spatial excitation, the gradientcoils 120 can attach spatially specific frequency and phase informationto the atomic nuclei placed within the imaging volume, allowing theresultant MR signal to be reconstructed into a useful image. A gradientcoil control unit 125 in communication with the data processing system105 can be used to control the operation of the gradient coils 120.

In some implementations of the DREMR system 100, there may be additionalelectromagnet coils present, such as shim coils (traditionally, but notlimited to, producing magnetic field profiles of 2nd order or higherspherical harmonics) or a uniform field offset coil or any othercorrective electromagnet. To perform active shimming (correcting thefield distortions that are introduced when different objects are placedwithin or around the system), the corrective electromagnets, such as theshim coils, carry a current that is used to provide magnetic fields thatact to make the main magnetic field more uniform. For example, thefields produced by these coils can aid in the correction ofinhomogeneities in the main magnetic field due to imperfections in themain magnet 110, or to the presence of external ferromagnetic objects,or due to susceptibility differences of materials within the imagingregion, or any other static or time-varying phenomena.

The DREMR system 100 can further comprise radio frequency (RF) coils130. The RF coils 130 are used to establish an RF magnetic field with astrength B1 to excite the atomic nuclei or “spins”. The RF coils 130 canalso detect signals emitted from the “relaxing” spins within the objectbeing imaged. Accordingly, the RF coils 130 can be in the form ofseparate transmit and receive coils or a combined transmit and receivecoil with a switching mechanism for switching between transmit andreceive modes.

The RF coils 130 can be implemented as surface coils, which aretypically receive only coils and/or volume coils which can be receiveand transmit coils. The RF coils 130 can be integrated in the main fieldmagnet 110 bore. Alternatively, the RF coils 130 can be implemented incloser proximity to the object to be scanned, such as a head, and cantake a shape that approximates the shape of the object, such as aclose-fitting helmet. An RF coil control unit 135 in communication withthe data processing system 100 can be used to control the operation ofthe RF coils 130.

To create a contrast image in accordance with field-shifting techniques,DREMR system 100 can use field-shifting electromagnets 140 whilegenerating and obtaining MR signals. The field-shifting electromagnets140 can modulate the strength of the main magnetic field. Accordingly,the field-shifting electromagnets 140 can act as auxiliary to the mainfield magnet 110 by producing a field-shifting magnetic field thataugments or perturbs the main magnetic field. A field-shiftingelectromagnet control unit 145 in communication with the data processingsystem 100 can be used to control the operation of the field-shiftingelectromagnets 140.

To reduce interference and artefacts, the field-shifting electromagnets140 may include a shield similar to the shielded gradient coils 120described above. The shielded field-shifting electromagnets 140 can havetwo components: an inner primary field-shifting electromagnets, toproduce the field shift and an outer shield field-shiftingelectromagnets, to form a shield by reducing the stray field of theprimary field-shifting electromagnets outside a certain volume such asan imaging volume. Implementing field-shifting primary and shieldelectromagnets combination that balances the competing needs of lowinductance (faster slew rates), high efficiency (greater magnetic fieldstrength for a given current amplitude), and low resistance (lessheating and subsequent demands on cooling) is a complex electromagneticproblem.

Indeed, one side effect of shielding the field-shifting electromagnets140 can be that the fields produced by the primary and shield componentsof the shielded field-shifting electromagnets 140 may partially canceleach other within the imaging volume. Accordingly, more current can berequired to produce a magnetic field with a particular strength byshielded field-shifting electromagnets 140 than by unshieldedfield-shifting electromagnets 140. This effect can be quantified as thefield-shift efficiency, which may be defined as the field-shiftamplitude per 1 Ampere of current passing through the electromagnet. Theloss in efficiency for a shielded configuration depends on the distanceand current density ratio between the shield electromagnets and theprimary electromagnets. Increasing the distance between the primary andshield electromagnets may increase the field-shift efficiency.

The conductive components of the field-shifting electromagnets 140,including the primary and shield electromagnets, may consist of anelectrical conductor (for example copper, aluminum, etc.). The internalelectrical connections can be such that when a voltage difference isapplied to the terminals of the field-shifting electromagnets 140,electric current can flow in the desired path. The conductive componentsfor both the primary and the shield electromagnets can be insulated byphysical separation and/or a non-conductive barrier. The field-shiftwindings can be placed in layers on or within a non-conductive substrate(for example, G10, FR4, epoxy or others).

In some variations, the field-shifting electromagnets 140 may also beprovided with thermal control or heat extraction mechanisms. Forexample, where windings are used to form the electromagnets, thewindings can be hollow and coolant can be passed through these hollowconductors to extract heat deposited in the electromagnet due toresistive heating of the windings when electricity is applied.Alternatively, other methods of extracting heat can be used, such asinserting coolant channels within the field-shifting electromagnets 140.The coolant channels can be in thermal contact with the field-shiftingelectromagnets 140. The field-shifting electromagnets 140 can also bemounted in a thermally-conductive but electrically-non-conductive epoxyto ensure that the mechanical assembly is rigid and to limit thepossibility of electrical breakdown.

There are many techniques for obtaining images using the DREMR system100, including T1 and T2 weighted images. To provide a simplifiedillustration of the DREMR system 100's functionality, simplifiedoperations for obtaining proton density-weighted images are described asa non-limiting example. To create an image in accordance with theexample illustration, the DREMR system 100 detects the presence ofatomic nuclei containing spin angular momentum in an object, such asthose of Hydrogen protons in water or fat found in tissues, bysubjecting the object to a relatively large magnetic field. In thisexample implementation, the main magnetic field has a strength of B0 andthe atomic nuclei containing spin angular momentum may be Hydrogenprotons or simply protons. The main magnetic field partially polarizesthe Hydrogen protons in the object placed in the imaging volume of themain magnet 110. The protons are then excited with appropriately tunedRF radiation, forming an RF magnetic field with a strength of B1, forexample. Finally, weak RF radiation signal from the excited protons isdetected as an MR signal, as the protons “relax” from the magneticinteraction. The frequency of the detected MR signal is proportional tothe magnetic field to which they are subjected. Cross-sections of theobject from which to obtain signals can be selected by producing amagnetic field gradient across the object so that magnetic field valuesof the main magnetic field can be varied along various locations in theobject. Given that the signal frequency is proportional to the variedmagnetic field created, the variations allow assigning a particularsignal frequency and phase to a location in the object. Accordingly,sufficient information can be found in the obtained MR signals toconstruct a map of the object in terms of proton presence, which is thebasis of a traditional MRI image. For example, since proton densityvaries with the type of tissue, tissue variations can be mapped as imagecontrast variations after the obtained signals are processed.

Referring now to FIG. 2, to further illustrate the example signalacquisition process by the DREMR system 100, it will be assumed that anobject is placed within an imaging volume 250 of the main magnet 110having a main magnetic field 210 with a strength B0, pointing along theZ-axis indicated at 240. The object subsequently has a net magnetizationvector. In this illustrative example, a slice in a plane along the X andY axes, as indicated at 205, is being imaged. It should be noted that inthis example, the slice has a finite thickness along the Z-axis,creating a volumetric slice 205.

When the object is placed in the main magnetic field B0, the individualspins align themselves in the direction of the Z-axis. Referring to FIG.3, at equilibrium, the magnetization by main field B0 can produce a netlongitudinal magnetization Mz, with an amplitude of M0, parallel withthe main magnetic field. Excitation of the spins may be achieved when aradio frequency (RF) pulse that generates the RF magnetic field with anamplitude of B1 is applied at the Larmor frequency, by the RF coils 130.During the application of the RF magnetic field the net magnetizationrotates around the applied RF (B1) field and can cause the netmagnetization to rotate away from the Z-axis. The component of therotated magnetization that is projected in the X-Y plane is the nettransverse magnetization Mxy. The spins can precess about the applied RFmagnetic field until the RF magnetic field is removed.

Once the equilibrium magnetization is perturbed, spin-relaxationprocesses occur. Spin-lattice relaxation processes cause a return ofmagnetization to the equilibrium distribution along the Z-axis.Spin-lattice relaxation can thus bring the longitudinal magnetization Mzback toward its maximum value M0, as indicated at 305, with acharacteristic time constant T1. A characteristic time representing therecovery of the magnetization along the Z-axis by 37% is called the T1relaxation time or T1 time. 1/T1 is referred to as the longitudinalrelaxation rate.

Spin-spin relaxation, on the other hand, can cause a loss of coherencedue to dephasing of the net transverse magnetization. Therefore, duringspin-spin relaxation, the transverse magnetization Mxy exponentiallydecays toward zero, as indicated at 310, with a characteristic timeconstant T2. A characteristic time representing the decay of the signalby 37%, is called the T2 relaxation time or T2 time. 1/T2 is referred toas the transverse relaxation rate.

Transverse relaxation (T2) can cause irreversible dephasing of thetransverse magnetization. There is also a reversible dephasing effectcaused by magnetic field inhomogeneities. These additional dephasingfields may arise from a variety of sources including the main magneticfield inhomogeneity, the differences in magnetic susceptibility amongvarious tissues or materials, chemical shift, and gradients applied forspatial encoding. The contribution to the transverse relaxation timefrom these reversible dephasing effects are typically referred to asT2′. The characteristic relaxation time of the combination of reversible(T2′) and irreversible (T2) dephasing effects is typically referred toas T2* relaxation.

The difference between the time constants T1 and T2 are important fordevelopment of contrast in MR imaging. The relaxation times can varywith the strength of the magnetic field applied, as well as temperature.Moreover, T1 and T2 values associated with biological tissues can vary.Generally, tissues with shorter T1 times, such as T1a as indicated at315, can yield greater signal intensity at a given point in time(appearing brighter in images) than those with longer T1 times, such asT1b as indicated at 305, due to the more rapid recovery of signal. Onthe other hand, tissues possessing short T2 times, such as T2a asindicated at 320, can yield lower signal intensity (appearing darker inimages) due to a reduction in the detected transverse magnetization Mxy.The MR signal from an image can be therefore dependent on thecombination of the intrinsic tissue properties and extrinsicuser-selected imaging parameters and contrast agents.

To obtain images from the DREMR system 100 in the traditional manner,one or more sets of RF pulses and gradient waveforms (collectivelycalled “pulse sequences”) are selected at the data processing system105. The data processing system 105 passes the selected pulse sequenceinformation to the RF control unit 135 and the gradient control unit125, which collectively generate the associated waveforms and timingsfor providing a sequence of pulses to perform a scan.

The sequence of RF pulses and gradient waveforms, namely the type ofpulse sequence, applied may change which relaxation times have the mostinfluence on the image characteristics. For example, T2* relaxation hasa significant influence following a 90° RF pulse which is used in agradient-echo (GRE) sequence, whereas T2 relaxation has a moresignificant influence following 90°-180° sequential RF pulses (alsoknown as a spin echo sequence).

Referring now to FIG. 4, an illustrative pulse sequence 400 is shownthat can be used to acquire images using the DREMR system 100.Specifically, a timing diagram for the example pulse sequence isindicated. The timing diagram shows pulse or signal magnitudes, as afunction of time, for transmitted (RFt) signal, magnetic field gradientsG_(x), G_(y), and G_(z), received RFx signal and filed-shifting signal(FS). An idealized pulse sequence, simplified for illustrative purposes,can contain a slice selection radio frequency pulse 410 at RFt, a sliceselection gradient pulse 420 at Gz, a phase encoding gradient pulse 430at Gy, a frequency encoding gradient pulse 440 at Gx, as well as adetected MR signal 450 at RFx. The pulses for the three gradients Gx,Gy, and Gz represent the magnitude and duration of the magnetic fieldgradients that can be generated by the gradient coils 120. The sliceselection pulse 410 can be generated by the transmit aspect of RF coils130. Detected MR signal 450 can be detected by the receive aspect of theRF coils 130. In this illustrative example it will be assumed thattransmit aspect and receive aspect of RF coils 130 are formed bydistinct coils. Finally, the field-shifting signal FS causes the mainmagnetic field strength to be changed for the duration of the signal FS.The precise timing, amplitude, shape and duration of the pulses orsignals may vary for different imaging techniques. For example,field-shifting signal FS may be applied at a time and manner that allowsimage contrast to increase for the technique used.

The first event to occur in pulse sequence 400 can be to turn on theslice selection gradient pulse 420. The slice selection RF pulse 410 canbe applied at the same time. In this illustrative example, the sliceselection RF pulse 410 can be a sinc function shaped burst of RF energy.In other implementations, other RF pulse shapes and durations can beused. Once the slice selection RF pulse 410 is turned off, the sliceselection gradient pulse 420 can also be turned off and a phase encodinggradient pulse 430 can be turned on. In some implementations, thefield-shifting signal 460 may also be turned on at this point to changethe main magnetic field strength. Once the phase encoding gradient pulse430 is turned off, a frequency encoding gradient pulse 440 can be turnedon and a detected MR signal 450 can be recorded. It should be noted thatthe shapes, magnitudes and durations of the pulses and signals shown inFIG. 4 are chosen for illustrative purposes, and that inimplementations, one or more of these factors and others may be variedto achieve the desired scan results.

The pulse sequence 400 can be repeated a certain number of times oriterations, typically 256 times, to collect all the data needed toproduce one image. The time between each repetition of the pulsesequence 400 can be referred to as the repetition time (TR). Moreover,the duration between the center point of the slice selection pulse 410and the peak of detected MR signal 450 can be referred to as echo time(TE). Both TR and TE can be varied as appropriate for a desired scan.

To further illustrate the signal acquisition process of DREMR system100, FIG. 2 is referred to in conjunction with FIG. 4. To select aslice, the slice selection gradient pulse 420 can be applied along theZ-axis, satisfying the resonance condition for the protons located inthe slice 205. Indeed, the location of the slice along the Z-axis can bedetermined based in part on the slice selective gradient pulse 420.Accordingly, the slice selection pulse 410, generated at the same timeas the slice selection gradient pulse 420 can excite protons that arelocated within the slice 205 in this example. Protons located above andbelow the slice 205 are typically not affected by the slice selectionpulse 410.

Continuing with the illustrative example, in accordance with the pulsesequence 400, a phase encoding gradient pulse 430 can be applied afterthe slice selection gradient pulse 420. Assuming this is applied alongthe Y-axis, the spins at different locations along the Y-axis can beginto precess at different Larmor frequencies. When the phase encodinggradient pulse 420 is turned off, the net magnetization vectors atdifferent locations can precess at the same rate, but possess differentphases. The phases can be determined by the duration and magnitude ofthe phase encoding gradient pulse 430.

Once the phase encoding gradient pulse 430 is turned off, a frequencyencoding gradient pulse 440 can be turned on. In this example thefrequency encoding gradient is in the X direction. The frequencyencoding gradient can cause protons in the selected slice to precess atrates dependent on their X location. Accordingly, different spatiallocations within the slice are now characterized by unique phase anglesand precessional frequencies. RF receive coils 130 can be used toreceive the detected signal 450 generated by the protons contained inthe object being scanned while the frequency encoding gradient pulse 440is turned on.

As the pulse sequence 400 is performed by DREMR system 100, the acquiredsignals can be stored in a temporary matrix referred to as k-space, asshown in FIG. 5 at 500. Typically, k-space is the collection of thedetected signals measured for a scan and is in the spatial frequencydomain. K-space can be covered by frequency encoding data along theX-axis 520 (Kx) and phase encoding data along the Y-axis 530 (Ky). Whenall the lines for the k-space matrix for a slice are received (at theend of the scan of a single slice, for example) the data can bemathematically processed, for example through a two-dimensionalFourier-transform, to produce a final image. Thus, k-space can hold rawdata before reconstruction of the image into the spatial domain.Typically, k-space has the same number of rows and columns as the finalimage and is filled with raw data during the scan, usually one line perpulse sequence 400. For example, the first line of k-space 500,indicated at 510, is filled after the completion of the first iterationof the pulse sequence generated for scanning a slice and contains thedetected signal for that pulse sequence iteration. After multipleiterations of the pulse sequence, the k-space can be filled. Eachiteration of the pulse sequence may be varied slightly, so that signalsfor the appropriate portions of the k-space are acquired. It should benoted that based on different pulse sequences, other methods of fillingthe k-space are possible, such as in a spiral manner, and arecontemplated.

The choice of specific pulse sequences with optimized parameters can beused by the DREMR system 100 to exploit tissue contrast to obtain imagesthat are able to depict different characteristics of tissue andmaterials. For example, as mentioned above, T2* relaxation has asignificant contribution on relative signal intensities immediatelyfollowing a 90° RF pulse. T2* relaxation can be one of the maindeterminants of image contrast with GRE pulse sequences and forms thebasis for many magnetic resonance (MR) applications, such assusceptibility-weighted imaging (SWI), perfusion MR imaging, andfunctional MR imaging. GRE sequences with T2*-based contrast can be usedto depict hemorrhage, calcification and iron deposition in varioustissues and lesions.

SWI uses phase information in addition to T2* relaxation based contrastto exploit the magnetic susceptibility differences of blood and of ironand calcification in various tissues. Accordingly, SWI is an MR imagingmethod that takes advantage of signal loss and phase information toallow better imaging of vessels and other tissues.

Functional MRI (fMRI) studies rely on regional differences in cerebralblood flow to delineate regional activity. Blood Oxygenation LevelDependent Imaging (BOLD) is a technique used to generate images infunction MRI studies. BOLD-fMRI is able to detect differences incerebral blood flow in part due to a difference in the paramagneticproperties of oxygenated hemoglobin and deoxygenated hemoglobin.Deoxygenated hemoglobin can be more strongly paramagnetic thanoxygenated hemoglobin, and therefore the former can cause greater localdephasing of protons. The local dephasing can reduce the MR signal fromthe tissues in its immediate vicinity. T2* weighted pulse sequences canbe used to detect this change.

The DREMR system 100 can also be used to perform MR spectroscopy.Spectroscopy is the determination of the chemical composition of asubstance by observing the spectrum of electromagnetic energy releasedfrom a material, including chemical samples, or a tissue sample. MRspectroscopy is a technique whereby MR signals obtained from the nucleiof a material is analyzed to identify the material's composition. MRspectroscopy is based on the fact that components of a material havedifferent resonant frequencies. Rather than displaying MR signals on agray scale as an image based on the relative signal strength, MRSpectroscopy displays the MR signal as a spectrum graph. Accordingly,the resonance frequency of each compound is represented on a graph as apeak.

MR spectroscopy can be performed with a variety of pulse sequences. Abasic sequence consists of a 90 degree RF pulse followed by reception ofthe MR signal by the receiving components of the RF coils 130, withoutany intervening gradient pulses. Moreover, many pulse sequences used forimaging, such as a spin echo sequence, can be used for MR spectroscopyas well.

A DREMR system 100 can enhance traditional MR images by modulating orvarying the strength B0 of the main magnetic field during at least aportion of one or more pulse sequences. To perform field-shifting scansusing a DREMR system 100, magnetic strength level B0 of the mainmagnetic field may be caused to rapidly, and uniformly change during oneor more portions of one or more pulse sequences used to obtain imagesignals which can form an image. The goal is to cause shifts in the mainfiled by a predetermined field-shifting magnetic field without causingartifacts or image degradation due to changes in the main magnetic field

Specifically, field-shifting electromagnets 140 can be used forobtaining a contrast image by causing a shift in the main magnetic fieldstrength. A field-shifting magnetic field can be applied during aportion of a pulse sequence causing the main magnetic field to be fieldshifted in strength. More specifically, the static magnetic fieldstrength B0 generated by the main magnet 110 can be either increased ordecreased by an amount dB through the use of field-shiftingelectromagnets 140. The field-shifting magnetic field generated by thefield shifting electromagnets 140 may be applied during part,substantially all, or all of a pulse sequence.

Field-shifting properties of DREMR system 100 can be combined withvarious traditional imaging techniques by modifying traditional pulsesequences as appropriate, and by including an appropriate field-shiftingsignal, to obtain improved images. For example, in certain types of MRimaging it is often desirable to suppress MR signals arising fromdifferent materials. A common example of this is the suppression of MRsignals arising from fat while preserving MR signals arising from water.This suppression can be done by making use of the fact that MR signalsfrom different materials may have different frequencies of precession.For example, protons of fat and water have different precessional orLarmor frequencies. Thus, in a homogenous main magnetic field, asufficiently narrow band RF pulse may be generated by RF coils 130 toexcite the desired tissue type only. If such a pulse is used to excitewater, for example, in place of the typical slice selection transmission410, it may primarily tip the magnetization of water molecules into thetransverse plane. Hence the resulting MR signal measured will primarilybe from the water molecules.

In alternative implementations, a saturation pulse may instead beapplied to suppress signals from the unwanted tissue types, such as fat.Thus, a sufficiently narrow band saturation pulse may be used by theDREMR system 100 to tip the protons of the undesired species into thetransverse plane. If such a pulse is used to suppress signals from fatprotons, for example, then a conventional slice select pulsecombination, such as pulse 410 and 420, applied shortly thereafter canprimarily tip the magnetization of water protons into the transverseplane since the fat protons would have already been excited by thesaturation pulse prior to the application of the slice selection pulse.Because the longitudinal magnetization of fat protons would not have hadtime to regrow, fat protons would not be available to tip into thetransverse plane at the time the slice selection pulse is applied. Thus,the resulting measured MR signal would be primarily obtained from thewater protons. The selective RF pulse used to excite the desired speciesmay be referred to as a saturation pulse.

One difficulty of the saturation method can be that the difference inprecessional frequencies between materials is proportional to the mainmagnetic field strength. At lower main magnetic field strengths, theseparation between the precessional frequencies of protons of differentmaterials is lower. For example where B0 is at 0.5 T, the separationbetween precessional frequencies of fat and water protons (whoseprecessional frequencies differ by 3.5 parts per million), isapproximately 70 Hz whereas at 1.5 T the separation is approximately 220Hz. FIG. 6(a) illustrates a generic 15 ms radio frequency saturationpulse response 605 for exciting water, compared to signals from fat(610) and water (615) at one hypothetical main magnetic field strengthB0 strength. As illustrated in FIG. 6(b), at a lower strength B0′ and asimilar duration saturation pulse, the saturation pulse response 605 isnot sufficient for robust saturation. It should be noted thatillustrations of FIG. 6 are not to scale and the elements have beenchosen to clarify the concepts being discussed.

Additional problems involve criteria used for generating narrow bandsaturation pulses. Saturation pulses which are designed to affect only anarrow range of frequencies are generated in accordance with variouspractical constraints including how sharply the frequency-dependenteffect can occur, how long the RF pulse takes, how much RF power isneeded and other criteria. Accordingly, generating effective narrow bandsaturation pulses get increasingly difficult as the Larmor frequencyseparation between tissue types decreases.

By applying a field-shifting magnetic field, generated for example, bythe field-shifting coils 140, the strength B0 of the main magnetic fieldcan be increased by dB during the spectral selective or saturationportion of the MR pulse sequence. Thus the separation between theprecessional frequencies of different materials can be increased,allowing the use of saturation pulses that are more practical andeffective. In accordance, a spectrally selective saturation pulse can bedesigned for a main field strength of B0+dB where dB is the strengthadded by the magnetic field generated by field-shifting coils 140.

Referring to FIG. 7, an example method of augmented MR signalacquisition is illustrated. A saturation pulse can be combined with apredetermined pulse sequence, such as pulse 400, to effect MR imageacquisition. Accordingly, at 705, the saturation portion of the combinedpulse sequence, the saturation pulse is generated by RF coils 130,concurrently with the field-shifting magnetic field, as generated byfield-shift coils 140 to increase the main magnetic field strength toB0+dB. The increase, in turn, allows a greater separation of theprecession frequencies of different materials, increasing the efficacyof the saturation pulse. After the saturation portion 705, thepredetermined portion 710 of the combined pulse sequence is applied.During the predetermined portion 710, the field-shifting field may beturned off and a predetermined pulse sequence such as that of pulsesequence 400 may be applied, the pulse sequence being designed for themain magnetic field strength B0. This process may then be repeated asshown at 715 and 720. The repetition may last as many times as desiredto obtain appropriate MR images. In variations, the spectral saturationportion of the combined pulse sequence may not always be provided priorto the beginning of the predetermined pulse sequence. In somevariations, the spectral saturation portion may be applied at some pointwithin the predetermined pulse sequence. In further variations, thefield-shifting field may also be applied during at least a portion ofthe predetermined portion 710 of the pulse sequence, the pulse sequenceapplied being appropriately varied to account for the shifted strengthof the main magnetic field. The additional application of thefield-shifting field during a pulse sequence portion may be at adifferent strengths, such as dB1, than the field-shifting field appliedduring a spectral saturation portion. Moreover, each repetition mayinvolve field-shifting fields that are different in strength andduration than the previous application of the field-shifting field.

Field-shifting properties of DREMR system 100 can also be combined withsusceptibility-weighted imaging (SWI). SWI is an MR imaging method whereimage contrast is generated based on local variations in the magneticfield caused by local magnetic susceptibility variations of materials.SWI uses phase information in addition to T2*-relaxation time basedcontrast to exploit the magnetic susceptibility differences of tissuesand/or materials such as blood and iron. In other words, SWI is animaging method where image contrast may be enhanced based on magneticsusceptibility differences between tissues and/or materials.

Magnetic susceptibility is a property of material which determines analteration in a magnetic field caused by a material, when that materialis placed in a magnetic field, such as the main magnetic field during MRimaging. For example, the magnetic field strength H inside a tissue,depends on that tissue's magnetic susceptibility which is an inherentproperty of the tissue. The relationship between the strength H of thesusceptibility altered magnetic field and the main magnetic field, B0,can be expressed as H=(1+x)*B0 where x is the magnetic susceptibilityproperty of the material. For example, venous blood has a xapproximately equal to −6.56×10⁻⁶ and soft tissues have a xapproximately equal to −9.05×10⁻⁶. Accordingly, SWI imaging can be usedto image the difference in susceptibility altered magnetic fieldsbetween venous blood and soft tissues as caused by susceptibilitydifference between the two tissue types.

As an example, venous blood and hemorrhage (bleeds) areas have asusceptibility difference from soft tissue. This difference can causethe venous blood, or hemorrhage areas, to have a signal with a shorterT2* in comparison with soft tissues. Accordingly, signals from venousblood/bleeds can decay away faster and produce less signal in aT2*-weighted pulse sequence (e.g. a GRE sequence).

The strength of the main magnetic field can be another factor thataffects the differences in susceptibility altered magnetic field betweentissues. Accordingly, increasing the magnetic field applied to an objectduring imaging through the application of a field-shifting magneticfield, can increase, for example, the imaged contrast between blood suchas venous blood and other tissues obtained by SWI imaging. For example,a typical SWI pulse sequence can be generated while the main magneticfield with a strength of B0 is supplemented by the field-shiftingmagnetic field generated by field electromagnets 140, increasing themain magnetic field strength to B0+dB. The field-shifting magnetic fieldmay be applied during the interval between signal excitation andacquisition.

Referring to FIG. 8, an illustrative example method for augmenting SWIwith the use of field-shifting magnetic field using the DREMR system 100is indicated. Excitation is achieved, through application of an RF pulseby the RF coils 130, at an excitation portion 810 of a SWI pulsesequence 805 for acquiring an SWI image. The main magnetic fieldstrength is at B0. At the phase accrual portion 815, of the SWI pulsesequence 805, which is the time during which much of themagnetic-susceptibility-based image contrast is generated, afield-shifting field is applied by the field-shifting coils 140, whichcauses the strength of the main magnetic field to be increased to B0+dBas indicated at 820. Next, the data acquisition portion 825 of the SWIpulse sequence allows acquisition of the MR signals. The process can berepeated, as indicated at the second SWI pulse sequence 830. Therepetition may occur a predetermined number of times to obtain a desiredimage. It should be noted that a field-shifting magnetic field may beapplied during portions of the pulse sequence other than the phaseaccrual portion and the pulse sequence portions adjusted as desired inaccordance with the changed main magnetic field strength. Moreover, thestrength and the duration of the field-shifting field applied may varyat different portions or different repetitions of the SWI pulsesequence.

In variations, SWI may be used as a method that can help visualize smallbleeds in tissue. In some situations, such as small tissue regions wherehemorrhaging has occurred or small areas of blood, detecting thecontrast difference due to susceptibility effects can be challenging,especially at lower magnetic field strengths where the susceptibilityeffect is reduced compared to high fields. In these situations, thereduced variation in signal strength due to the susceptibility effectmay be enhanced by combining images with different levels ofsusceptibility weighting. This can be achieved by acquiring images atdifferent main magnetic field strengths. As an example, for sometissues, the corresponding signal obtained in an SWI image can be highbut may not change significantly when the image is acquired usingdifferent main magnetic field strengths. Furthermore, there may also bea region of a small bleed (background tissue into which blood hashemorrhaged) for which the corresponding SWI image signal can be low butmay change significantly with different main magnetic field strengths.If the small bleed region is embedded within the background tissue, theimage contrast between an image location containing background tissueonly and an image location containing a region of small bleed would beproportionally small. If two images are acquired at two differentmagnetic field strengths and the images are subsequently subtracted, thebackground tissue signal would be eliminated and the relative contrastbetween the region containing background tissue only and one containingbackground tissue and a small bleed would be increased.

As an illustrative example, an SWI image can be acquired in accordancewith a SWI pulse sequence at a first main magnetic field strength, suchas B0. The acquisition can be followed by the acquisition of one or moreadditional susceptibility weighted images using the same SWI pulsesequence, but at different main magnetic field strengths as achievedthrough the application of a field-shifting magnetic field byfield-shifting coils 140. The images from each of these acquisitions,each image being acquired at a different main magnetic field strength,can then be combined to produce an image that emphasizes regions wherethe susceptibility-induced contrast varied from image to image based onthe variation of the main field strength to field strength. The twoimages may be combined in any manner that can increase the relativeimage contrast. This may include subtracting images in pairs; summingall the images together; fitting the signal at each pixel locationacross the images to some parametric model; or other mathematicalcombinations.

Referring to FIG. 9, a simplified example of a method for visualizingsmall bleeds in tissues using the DREMR system 100 is illustrated.Excitation can be achieved, through application of an RF pulse by the RFcoils 130, at excitation portion 910 of a SWI pulse sequence 905 foracquiring an SWI image. At the phase accrual portion 915, of the SWIpulse sequence 905, which is typically the time during which much of themagnetic-susceptibility-based image contrast can be generated, afield-shifting field can be applied by the field-shifting coils 140, atleast during a part of the portion 915. The application of thefield-shifting field typically causes the strength of the main magneticfield to be increased to B0+dB as indicated at 920. Next, the dataacquisition portion 925 of the SWI pulse sequence can allow theacquisition of the MR signals and thus a portion of an MR image. Theprocess can then be repeated, as indicated at the second SWI pulsesequence 930. However, during the SWI pulse sequence 930, thefield-shifting field applied by the field-shifting coils 140 asindicated at 935 is at a strength dB1, different from the initialapplication of the auxiliary field at strength dB indicated at 920. Itshould be noted that pulse sequence 930 is typically the same pulsesequence as pulse sequence 905, altered as necessary to accommodate thechanges in the main magnetic field. The variations in main fieldstrength to dB1 and dB can coincide in location and duration within thetwo pulse sequences. The pulse sequence pairs may be repeated, apredetermined number of times, as they are varied appropriately toobtain two complete images. In variations, the two images may beacquired sequentially. For example, a number of pulse sequences desiredto obtain a first image may be applied at a first main magnetic field,and repeated at a second main magnetic field strength to obtain a secondimage. In other variations, other methods for obtaining two images attwo different main magnetic field strengths can be used. To generate thefinal contrast enhanced image, the two images can be combined asdescribed above. It should be noted that a field-shifting magnetic fieldmay be applied during portions of the scan other than the phase accrualportion. For example, the field-shifting magnetic field can remain onduring data acquisition, or for part of the data acquisition. In furtherimplementations, the strength of the field-shifting field applied mayvary within or at different portions of the SWI pulse sequence.

The process of acquiring multiple images at differing main magneticfields field-shifted by field-shifting coils 140 may be repeated as manytimes as required. For example, in some implementations, more than twoimages may be acquired. When more than two images are used they may becombined in any manner that can increase the relative image contrast.This may include subtracting images in pairs, then summing thesubtracted images; summing all the images together; fitting the signalat each pixel location across all the images to some parametric model;or other mathematical combinations. In further implementations, thestrength of the field-shifting field applied may vary within or atdifferent portions of the SWI pulse sequence. For example, the auxiliaryfiled can remain on during the data acquisition, or for part of the dataacquisition.

Field-shifting properties of DREMR system 100 can also be combined withother T2*-weighted MR imaging techniques. As discussed above, T2*relaxation refers to the decay of transverse magnetization caused by acombination of spin-spin relaxation and magnetic field inhomogeneity.T2* relaxation has contributions both from the T2 relaxation which is aninherent tissue property, as well as contributions from local magneticfield inhomogeneities, commonly referred to as the decay time T2′. Thethree relaxations are related by 1/T₂*1/T₂+1/T₂′, where T₂′≅γΔB₀ whereΔB0 measures the magnetic field inhomogeneities. Accordingly, T2*relaxation, as described above, can be detected with gradient-echo (GRE)imaging because transverse relaxation T2′ caused by magnetic fieldinhomogeneities, unlike in the case of a 180° pulse at spin-echoimaging, is not eliminated by a GRE pulse.

There can be many contributions to the magnetic field inhomogeneitiesincluding inhomogeneities in the main magnetic field due tocharacteristics of main magnet 110, as well as magnetic susceptibilitybased field differences. Both of these effects scale linearly with thestrength of the main magnetic field. Thus, the rate of signal decay T2′,and hence T2*, may vary in different materials placed within differentmain magnetic fields.

One or more T2*-weighted MR images may be acquired using knownT2*-weighted imaging methods, with a field-shifting magnetic field beingprovided by the field-shifting coils 140 for at least some of the imagesduring all or part of the time during which T2* decay occurs in thepulse sequence. The T2* dispersion signal can then be generated byobserving the variation in T2*-weighted signal at each magnetic fieldstrength for a given material such as tissue and/or region of the image,for example. Accordingly, changes in main field strength of DREMR system100 can be provided through variations in the field-shifting magneticfield applied by the filed-shifting coils 140. The variation of T2*dispersion signal in accordance with the main magnetic field can then beanalyzed to differentiate different tissues by identifying, for example,unique patterns in the relationship between T2* and magnetic fieldstrength or, as another example application, determine iron contentwithin the tissues. As further example, the T2* dispersion analysis caninclude the identification of unique magnetic field strengths wherethere is a rapid increase or decrease in the T2* dispersion curve thatmay be a unique characteristic for a given tissue.

Referring to FIG. 10, an example method of generating a T2* dispersionsignal using DREMR system 100 is illustrated. Excitation is achieved,through application of an RF pulse by the RF coils 130, at excitationportion 1010 of a T2* pulse sequence 1005 for acquiring T2* signal. Atthe T2* decay portion 1015, of the T2* pulse sequence 1005, afield-shifting magnetic field is applied by the field-shifting coils 140as indicated at 1020. Next, the data acquisition portion 1025 of the T2*pulse sequence allows acquisition of the MR signals. The process is thenrepeated, as indicated at the second T2* pulse sequence 1030 and thirdT2* pulse sequence 1035. However, during the second pulse sequence 1030,and the third pulse sequence 1035 the field-shifting magnetic fieldapplied by the field-shifting coils 140 as indicated at 1040 and 1045respectively is at strengths differing from the initial application ofthe field-shifting field indicated at 1020. Specifically, at 1040, themain magnetic field strength has been shifted to B0+dB1 and at 1045, themain filed strength has been shifted to B0+dB2. The repetition may occuran additional predetermined number of times. It should be noted that afield-shifting magnetic field may be applied during portions of thepulse sequence other than the T2* decay portion. In someimplementations, the strength and/or duration of the field-shiftingfield applied may vary within or at different portions of the T2* pulsesequence. For example, the field-shifting filed can remain on during orpart of the data acquisition portion of a pulse sequence. Although thisexample discusses obtaining and comparing signals associated with asingle pulse sequence repeated at different main magnetic fieldstrengths, the same process can be applied to entire images or portionsor regions of images acquired in a similar manner, using different mainfield strengths.

Once the multitude of signals or images are acquired at different mainfield strengths, they may be compared to determine changes in T2*dispersion. Referring to FIG. 11, a conceptual illustration of how T2*signals from 2 materials, indicated by a circle and a star, which couldbe the same (P1) at one field strength (B0+dB indicated at 1105 andcorresponding to signals acquired using T2* pulse 1005 of FIG. 10) canbe differentiated by repeating MR signal acquisition at shifted mainmagnetic fields. At magnetic field strength B0+dB1, indicated at 1110and corresponding to signals acquired using T2* pulse 1030 of FIG. 10,the T2* signals for the two materials are now different (P2 and P4). Atmagnetic field strength B0+dB2, indicated at 1115 and corresponding tosignals acquired using T2* pulse 1035 of FIG. 10, the T2* signals forthe two materials or tissues are further differentiated (P3 and P5).Based on these differentiations, the type of material can be determined.For example, the differentiation may simply indicate a specific magneticfield strength (which may be different from the unshifted main fieldstrength of the MRI system) at which there is the largest difference inT2* values between two tissues and at which T2* based imaging would bepreferably performed. Alternatively, the dispersion patterns for any setof tissues may suggest specific data processing to increase T2* basedsignal differentiation from the tissues. This could include fitting themeasured T2* dispersion points to a specific model (shape of variation),subtraction or other linear combinations of signals or images atspecific magnetic field strengths or other image combination methods.

As discussed above, T2′ component of T2*, and accordingly, T2* varieswith the applied magnetic field strength. For most materials or tissues,the expected variation of T2* with respect to main magnetic fieldstrength is linear. Specifically, the T2* change caused by an increasein the main magnetic field strength may be balanced by a T2* changecaused by a decrease in the field strength by the same amount. For somematerials, in particular those containing iron, the variation of T2*with respect to the field strength can be non-linear. The DREMR system100 can be used to take advantage of this non-linearity to performenhanced iron or BOLD imaging. T2* weighted images, both with andwithout main field perturbations, can be acquired. Such pairs of imagesmay be performed such that they differ in regions where the T2* responseto field variations is non-linear. For example regions containingiron-based compounds may show changes in contrast.

To implement a differential acquisition, a first acquisition may beperformed where no main field perturbation is applied. In a secondacquisition of the same MR image, the main field strength can be variedin a manner which can alter the image contrast for materials having anon-linear response to the field variation. As an example, the mainfield may be changed in one direction during a T2* decay portion of aT2* pulse sequence, and may be changed in an equal but oppositedirection, and for equivalent duration, for another portion of theT2*decay. For materials having a non-linear response to main magneticfield variations, the change in T2* dispersion when the main magneticfield increases by a predetermined amount may not be balanced by thechange in T2* dispersion when the main magnetic field decreases by anequal amount and duration. This may be in contrast to tissues ormaterials that vary linearly with respect to changes in the mainmagnetic field where the change in T2* dispersion can be the same whenthe main magnetic field is perturbed up and down by the same amount andduration.

Referring to FIG. 12, an example of a method for performing Iron or BOLDimaging using a DREMR system 100 is illustrated. In this figure pulsesequence 1205 is used to perform a T2*-weighted acquisition at mainfield strength B0, without any main field perturbations. Following theMR signal acquisition based on the pulse sequence 1205, the same pulsesequence is repeated at 1210. This time, however, the main fieldstrength B0 is increased by dB for a period of time during which T2*decay is occurring, through an application of a field-shifting magneticfield by filed-shifting coils 140. Following the increase, the mainfield strength is decreased by the same amount dB, again through theapplication of an auxiliary magnetic field by field-shifting coils 140for an equivalent duration. Subsequently, the MR data is acquired. Itshould be noted that although in this example, the main magnetic fieldstrength is first increased, and then decreased by the same amount foran equal duration, many different ways of perturbing the main magneticfield during the T2* decay portion of a pulse sequence is possible aslong as the perturbations occur in a manner which can alter the imagecontrast for materials having a non-linear response to field variations.For example, the main magnetic field may be altered in a manner suchthat the alterations are balanced. There are various methods forachieving balanced alterations. For example, in some variations a seriesof increases and decreases of equivalent amounts in the main magneticfield strength may be applied during the T2* decay portion of the pulsesequence. The main magnetic field may be increased first, then decreasedby an equivalent amount and duration, increased back up, and decreasedagain by an equivalent amount and duration to the last increase. Eachincrease-decrease pair may be by a different amount and duration.Moreover, the order of increase and decrease may change, and pairs maynot be located immediately adjacent to each other. Although this examplediscusses obtaining signals associated with a single pulse sequencerepeated at different main magnetic field strengths, it is to beunderstood that a similar process can be applied to the acquisition andanalysis of two or more images.

Field-shifting properties of DREMR system 100 can also be combined withMR spectroscopy. As discussed above, MR spectroscopy is a method wherebyMR data is acquired and processed to identify components of a substancethat have different resonant frequencies. The difference in resonantfrequencies may arise, for example, based on protons existing indifferent chemical environments within a compound or within differentcompounds within a material such as a tissue. MR spectroscopy is oftenused to analyze substances that are at a very low concentration and thusgenerate very low MR signals. Accordingly, a distribution of peaks atdifferent frequencies are developed from MR signals to identifydifferent tissues or materials. However, MR signals acquired alsoinclude significant noise. The noise is generally uniformly distributedacross all frequencies. Due to the low concentration and low signal ofcompounds in tissues or materials, it can be difficult to identify peaksabove random noise associated with signal acquisition. To counter thelow signal-to-noise ratio that is typical in these measurements, oftenMR signal acquisition is repeated and averaged together as a method ofaveraging down white noise.

The DREMR system 100 can allow performing improved MR spectroscopythrough acquisition of multiple MR signals at different main magneticfield strengths. The relative separation of peaks may be proportional tothe main magnetic field strength. Moreover, the size of the peaks mayalso weakly depend on the main magnetic field strength. On the otherhand, random noise signals do not typically vary, in a determinatemanner, with changes in the main magnetic field strength. Accordingly,MR signal acquisitions can be repeated using the DREMR system 100, withat least some of the repeats being made at differing main magnetic fieldstrengths. Any peaks present in the acquired MR signal may thus move byknown amounts based on the known shift in the main magnetic field.Accordingly, the acquired MR signals may be processed to identify peaksthat have shifted by the predicted amounts making it possible to improvethe detection of desired signal peaks.

Referring to FIG. 13, an example method for performing MR spectroscopyusing the DREMR system 100 is illustrated. FIG. 13(a) illustrates 3possible ideal signal peaks which are located at different frequenciesfor a given main field B0. FIG. 13(b) illustrates a characteristicrandom noise that is uniformly distributed over all frequencies and FIG.13(c) illustrates a combined signal resulting from the combination ofthe ideal peaks with the characteristic noise. FIG. 13(c) is indicativeof the type of signal that would be acquired by DREMR system 100.

Continuing with the figure, FIG. 13(d) illustrates the 3 possible idealsignal peaks corresponding to the three signal peaks of FIG. 13(a).However, in FIG. 13(d), the main magnetic field provided for acquisitionof the MR signals is increased by a field-shifting magnetic field with astrength dB. Accordingly, the position and amplitude of all 3 peaks arescaled relative to the 3 peaks of FIG. 13(a), in proportion to thechange of the main field strength from B0 to B0+dB. For example if B0+dBis equivalent to a main field with a strength of 2*B0, the position ofthe peaks in FIG. 13(d) may be scaled in frequency by a multiple of 2relative to the peaks in FIG. 13(a). FIG. 13(e) illustrates acharacteristic random noise acquired at the augmented main magneticfield strength B0+dB. As shown, the noise may not change with respect tothe change in the main field strength in a predictable and systematicway. Finally, FIG. 13(f) illustrates a combined signal resulting fromthe combination of the ideal peaks of FIG. 13(d) with the characteristicnoise of FIG. 13(e). FIG. 13(f) is indicative of the type of signal thatwould be acquired by DREMR system 100.

Correlating the peak movements with a change in the strength of the mainmagnetic field may help filter out randomly distributed noise which isinvariant with respect to changes in the main magnetic field strength.For example, to determine whether a peak is present in a pair of MRsignals acquired at different main magnetic field strengths, the peak'sexpected locations, which would be different at different main magneticfield strengths, can be checked in both MR signals. A peak can beassumed to be present when it is found in the expected locations in bothimages, the locations determined in part on the basis of the differencein main field strengths.

Although only two acquisitions were used in this illustrative example,additional acquisitions can be performed, and spectra obtained can beused in the determination of peak presence through the correlation ofpeak locations in the additional MR signals. In variations, each signalacquisition at a given main magnetic field strength may also be repeatedat the same magnetic field strength, and the signals thus acquiredaveraged to partially average out white noise as described above.

Field-shifting properties of DREMR system 100 can also be combined withMR fingerprinting. Any given tissue or material may be characterizedbased on a set of measured MR signal properties, referred to as the MRfingerprint of that tissue. For example, for a given tissue or material,multiple MR signal properties can be quantified on the basis of MRsignals acquired for that tissue or material. Accordingly T1, T2, T2*and/or other MR signal properties can be obtained for each tissue ormaterial based on MR signals acquired using one or more pulse sequences.These obtained set of MR signal parameters can then be used tocharacterize the MR scanned tissue or material.

MR signal properties can be dependent on the strength of the mainmagnetic field applied during signal acquisition. Accordingly, obtainingMR signal properties at multiple field strengths can add an additionaldimension to the set of parameters that could be used to characterizeand differentiate tissues. Accordingly, a set of MR signal propertiesare selected, and scans are performed with the DREMR system 100 usingappropriate pulse sequences for the selected MR properties to obtain theselected MR signal properties. The acquisition of the MR signalproperties are then repeated with different main magnetic fieldstrengths. The change in main magnetic field strength is accomplished byapplying an auxiliary magnetic field using the field-shifting coils 140.The field-shifting magnetic field can be can be specific to each MRsignal acquisition and applied in a manner that make MR signalmeasurements sensitive to changes in the main magnetic field strength.Some of these techniques, for example for the acquisition of T2*property, are discussed above. As a further illustrative example, toobtain a T1 measurement, well-established MR acquisition methods formapping the T1 relaxation parameter can be used. The acquisition can berepeated with an auxiliary field applied by field-shifting coils 140during the inversion time (TI) portion of the pulse sequence.

FIG. 14 provides a simplified illustrated example of how the addition ofmagnetic field strength changes can enhance MR fingerprinting. FIG.14(a) shows a distribution of values of one MR signal property,Parameter1, for two different tissue types at main magnetic fieldstrength B0 along the x-Axis. The y-axis has been added for illustrativeconvenience, and does not represent any values. According to FIG. 14(a),if the measured MR signal property Parameter1 falls between 1405 and1410, the tissue can be identified as Tissue A. If the measured MRsignal property Parameter1 falls between 1415 and 1420, on the otherhand, the tissue can be identified as Tissue B. The distribution ofvalues for the two tissues overlap. Accordingly, for a tissue ofinterest, after scanning the tissue, if a MR signal Parameter value of“Val1” was obtained, it would not be possible to uniquely identify whattissue type that value represented.

Continuing with the figure, FIG. 14(b) illustrates, along the x-axis, adistribution of values of one MR signal property (“Parameter 1”) for twodifferent tissue types at main magnetic field strength B0 along thex-Axis, as in FIG. 14(a). However, in this case, the y-Axis representsthe acquisition distribution of values of the same MR signal property,Parameter1, for the same two tissue types at main magnetic fieldstrength B0+dB. It can be noted that when the MR signal propertyParameter1 varies with respect to changes in the main magnetic fieldstrength, the measured MR signal property value for Parameter1 for agiven tissue will also change. Accordingly, the tissue of interestdiscussed above that had a Parameter1 value of Val1 at magnetic fieldstrength B0, may have a Parameter1 value of Val2 or Val3 atmagnetic-field strength B0+dB. Accordingly, if a tissue provides themeasurement Vail when measuring Parameter1 at B0 and Val2 when measuringParameter1 B0+dB, the tissue may be uniquely identified. Similarly, ameasurement of Val3 for Parameter1 at B0+db may uniquely identify thetissue of interest as being a different tissue. Note that this exampleis for 1 measured parameter whereas in practice more than one parametersmay be used, at least some which being magnetic field dependent. Theunique separation of tissue types may be resolved across multipleparameter dimensions, some of which will include one or moremeasurements at varying magnetic field strengths.

The addition of field-dependent contrast agents to tissues being scannedby DREMR system 100 pulse sequences can further enhance the detection oftraumatic brain injury (TBI), which is quite difficult to image usingtraditional MRI techniques. The contrast agents used typically have arelaxation profile that varies with the main field strength, both intheir bound and unbound state.

In some cases, an opening in the normally closed blood brain barrier(BBB) may allow albumin and fibrinogen to enter (normally inaccessible)brain tissue, and can be a specific cause of brain inflammation.Selectively imaging these molecules would enable imaging sites ofalbumin to and fibrinogen penetration in the brain, which may be used toidentify areas of brain trauma. To determine sites of increasedalbumin/fibrinogen in the brain, the patient can be injected with abolus of (appropriate) contrast agent and imaged with a field varyinggradient echo scan. By varying the field strength, molecular contrastcan be achieved. It would be advantageous to observe the quantity ofalbumin/fibrin in the brain, as well as to observe the time course ofthe spread of albumin/fibrin in a manner similar to current perfusionimaging.

The above-described embodiments are intended to be examples andalterations and modifications may be effected thereto, by those of skillin the art, without departing from the scope which is defined solely bythe claims appended hereto. For example, methods, systems andembodiments discussed can be varied and combined, in full or in part.

We claim:
 1. A delta-relaxation enhanced magnetic resonance (MR) imaging(DREMR) system comprising: a main magnet configured to generate a mainmagnetic field with a strength of B0; radio frequency coils having atransmit aspect and gradient coils configured to generate an initialpulse sequence for acquiring susceptibility weighted imaging (SWI)signals; field-shifting magnets configured to vary the main magneticfield strength to a strength of B1 during at least one portion of theinitial pulse sequence; and the radio frequency coils having a receiveaspect configured to acquire a first image based on the initial pulsesequence, wherein the initial pulse sequence is the pulse sequence foracquiring SWI signals, wherein the initial pulse sequence includes atleast one phase accrual portion and wherein the portion of the initialpulse sequence during which the main magnetic field strength is variedto B1 is at least a portion of the at least one phase accrual portion,wherein the transmit aspect of the radio frequency coils and thegradient coils further configured to generate a repeat pulse sequencefor acquiring SWI signals, the repeat pulse sequence corresponding tothe initial pulse sequence, the field-shifting magnets furtherconfigured to vary, for each portion of the initial pulse sequence wherethe main magnetic field strength is varied to B1, the main magneticfield strength of the corresponding portion of the repeat pulsesequence, to a strength of B2 different from B1, where the variation instrengths between B1 and B2 are not equal and opposite to each other,and the receive aspect of the radio frequency coils further configuredto acquire a second image based on the repeat pulse sequence, the DREMRsystem further comprising: a processing system configured to operateeach of the main magnet, the radiofrequency coils, the gradient coilsand the field-shifting magnets, and being further configured to combinethe first and the second images to produce a combined image emphasizingsusceptibility induced contrast.
 2. The DREMR system of claim 1 whereinthe processing system operates to combine by at least one of:subtracting one image from the other; summing the two images; andfitting the signal at each pixel location across the two images to aparametric model.
 3. The DREMR system of claim 1 wherein the initialpulse sequence is the pulse sequence for acquiring T2*-weighted MRimaging signals, wherein the initial pulse sequence includes at leastone T2* decay portion and wherein the portion of the pulse sequenceduring which the main magnetic field strength is varied to B1 is atleast a portion of the at least one T2* decay portion.
 4. The DREMRsystem of claim 1 wherein the initial pulse sequence is the pulsesequence for acquiring T2*-weighted MR imaging signals, wherein theinitial pulse sequence includes at least one T2* decay portion andwherein the portion of the pulse sequence during which the main magneticfield strength is varied to B1 is at least a portion of the at least oneT2* decay portion, wherein: the transmit aspect of the radio frequencycoils and the gradient coils further configured to generate a repeatpulse sequence for acquiring T2*-weighted MR imaging signals, the repeatpulse sequence corresponding to the initial pulse sequence; and eachportion of the repeat pulse sequence corresponding to the at least oneportion of the initial pulse sequence where the main magnetic fieldstrength is varied to B1, having a main magnetic field strengthdifferent from B1.
 5. The DREMR system of claim 1 wherein the initialpulse sequence is for acquiring saturation imaging signals, wherein theinitial pulse sequence includes at least one spectral saturation pulseand wherein the at least one portion of the initial pulse sequenceduring which the main magnetic field strength is varied to B1 is atleast a portion of the at least one spectral saturation pulse.
 6. TheDREMR system of claim 5 wherein the spectral saturation pulse suppressesMR signals from fat.